1. Field of the Invention
This invention relates in general to magnetic resonance imaging (MRI), and in particular to apparatus and methods for automatically tuning an MRI coil.
2. Description of Related Art
In an MRI process, a sample, such as a human being, is placed in a large magnetic field (the B.sub.0 field) that remains constant throughout the MRI process. The magnetic moment of nuclei in the body, in particular nuclei of hydrogen, become aligned with the magnetic field. Next the sample is exposed to an oscillating magnetic field having a selected frequency in the radio frequency (rf) region of the electromagnetic spectrum, causing the nuclei in the sample to resonate. The rf radiation is then switched off, but the nuclei continue to resonate resulting in the emission of rf radiation from the resonating nuclei. The emission is detected as an MRI signal
The resonance frequency of the sample depends upon the strength of the large magnetic field. This frequency is called the Larmor frequency and is expressed by the relationship L=gH, where L is the Larmor frequency, H is the strength of the magnetic field, g is a constant dependent upon the particular nuclei. The oscillating rf field is generated by an rf coil that encloses the sample. The frequency of the applied oscillating rf field is chosen to be substantially the same as the Larmor frequency.
The rf coil may also be used to receive the resonating emissions from the sample. The inductance and capacitance of the rf coil determine the tuned frequency of the coil and the impedance of the coil. The coil impedance is matched to the optimum source impedance of the preamplifier so that the noise figure of the preamplifier is minimized. The rf coil has its maximum sensitivity for the detecting emitted rf radiation when the inductance and capacitance of the rf coil are chosen so that the rf coil has a tuned frequency which is the same as the Larmor frequency of emitted rf radiation.
Prior art systems have, under certain conditions, had problems tuning the frequency of the rf coil to the Larmor frequency of the nuclei. Since the coil includes reactive elements (coil inductance and capacitive elements) the impedance of the coil is frequency dependent. Coil impedance has real and imaginary components that vary with frequency. Most coils have their impedance tuned and matched to maximize the signal-to-noise ratio of the detected signal. Often it is sufficient to adjust tuning and matching capacitors upon construction of the rf coil to match the source impedance of the preamplifier which minimizes the preamplifier's noise figure.
For an rf receive coil with a fixed geometry, the signal-to-noise ratio of magnetic resonance signals from a sample increases approximately linearly with the magnetic field. The closer the rf receive coil is to the sample, the larger the signal and the signal-to-noise ratio. Thus, for low fields it is very important that the receive coil be close to the body. The greater the distance between the coil and body, the poorer the MRI image.
The resonance frequency of a coil is determined by the reactive elements of the coil. The subject inside the coil increases it resistance which primarily affects the coil's bandwidth. The closer the coil is to the body, the larger the capacitive coupling to the body and the greater the shift in coil resonance frequency due to variances in the capacitive coupling between the coil and the body. A coil having a variable geometry will have a variable inductance. A change in coil inductance divided by the nominal coil inductance which is of the order of the Q of the loaded coil will severely detune the coil, resulting in reduced image quality. The resonance frequency of the coil may vary greatly from patient to patient due to changes in coil geometry from patient to patient.
Another factor affecting resonance frequency is variability in the value of the main magnetic field. Superconducting magnets produce high, stable magnetic fields (typically greater than 0.3T). Permanent magnets or conventional electromagnets can produce fields which often vary with time due to temperature variations, resulting in changes in the resonance frequency. Drifts in the magnetic field for a permanent magnet can be of the order of a thousand parts-per-million (PPM) per degree Centigrade resulting in a shift in the Larmor frequency. Such drifts result in the coil being severely detuned with respect to the Larmor frequency of the MRI system, leading to poor signal-to-noise ratios and, hence, poor image quality.
There are a number of techniques for tuning coils to the resonance frequency of the MRI system. See, for example, U.S. Pat. No. 4,897,604 which shows an expandable rf coil composed of a two parts including a main section and a removable bridge segment. Different size bridge segments change both the active and the physical circumference of the coil. In U.S. Pat. No. 5,143,068 a flexible coil having a fixed physical size is tuned by an externally located coupling coil circuit that has variable capacitors C.sub.s and C.sub.p. U.S. Pat. No. 4,791,372 also relies upon variable capacitors C.sub.s and C.sub.p.
Prior art systems have several limitations. For example, manually retuning an MRI coil for each patient is time-consuming and inefficient. The larger the frequency range over which tuning is required, the greater the time required, the more the difficulty in achieving a match, and thus the greater the cost. Imaging must be delayed until tuning is complete. Such delays add an extra financial burden to health care cost since the time of the patient and the MRI operator are taken up with the tuning process.
In the current art a single coil is not well suited for use with different sized patients because the frequency range for coil tuning and the coil filling factor cannot be made optimal for all patients. Prior art attempts to overcome these limitations with a multiplicity of discrete elements suffer from unreliability due to the number of connectors required between segments and are cumbersome to use since the coil must be reconstructed for each patient. See U.S. Pat. No. 4,897,604. Furthermore, each connection can add a series resistance to the coil due to contact resistance, resulting in a degradation of the coil signal to noise.